Developed in the 1970's,2, MRI has rapidly grown into an indispensable and increasingly popular imaging modality. Owing to its deep tissue penetration, non-invasiveness, excellent soft tissue contrast and high spatial resolution, MRI is used for diagnosis and treatment monitoring of a wide variety of diseases.
In conventional MRI scans, signals are mainly derived from 1H-NMR peaks of water and fat molecules present in the body being imaged. The image contrast of the tissues is determined by a number of factors, such as proton density, spin-lattice relaxation time (T1), and the spin-spin relaxation time (T2). T1 is a measure of how quickly the longitudinal magnetization vector (Mz) of spinning nuclei recovers towards the equilibrium direction after a resonant radio frequency (RF) pulse6. T2 relaxation time is a time constant that describes the dephasing of the transverse nuclear magnetization. Since T1 and T2 relaxations vary from tissue to tissue, acquisition parameters can be adjusted to differentiate among tissues. For example, dipoles in fats or hydrocarbon rich environments have much shorter T1 relaxation times than those in aqueous environments.
Despite its increasing number of applications, MRI is impeded by its intrinsic low detection sensitivity7,8 compared to imaging modalities that use ionizing radiation such as Positron Emission Tomography (PET), hampering its ability to detect certain pathologies, such as small tumors or differentiating post therapy tumor progression. Governed by the thermal equilibrium polarization of the nuclei, e.g. at room temperature and magnetic field of 1.5 Tesla (T) (commonly used in most clinical MR scanners), only 5 out of 1 million 1H spins are polarized9. According to Curie's Law, macroscopic magnetization is directly proportional to the magnetic field strength10. Increasing the field strength can partly compensates for this loss in sensitivity and improve signal-to-noise ratio (S/N)11. However, other than the cost of ultrahigh field scanners (higher than 7 T), there is a major concern relating to tissue overheating due to overexposure of radio frequency12 and technical issues such as coil design13.
Currently, the widely applied method of increasing MRI S/N, hence contrast and specificity, is the use of relaxation contrast agents, which can accelerate the relaxation rate of surrounding waters' nuclei spins. MRI CAs are categorized into T1 and T2 agents. T1 agents are mainly based on paramagnetic metal ions with unpaired valence electrons which can effectively shorten mainly the T1 relaxation time of the nearby water nuclei via electron-nuclear spin-spin coupling6. Clinical T1 CAs predominantly utilize Gd(III) which is chelated by different ligands to reduce the toxicity of free Gd(III) in vivo. Typical ligands include diethylene-triamine-penta-acetic acid (DTPA; Gd-DTPA is sold under the name Magnevist®) and 1,4,7,10-tetraazacyclododecane-1,4,7,10-tetraacetic acid (DOTA; Gd-DOTA is sold under the name Dotarem®) and their derivatives. T2 agents, which are mostly superparamagnetic particles, disrupt the homogeneity of the magnetic field causing predominant decreases in T2 and T2* due to diffusion of water through field gradients. T1 agents are able to generate positive contrast (increased signal intensity) in T1 weighted images while T2 agents such as superparamagnetic iron oxide nanoparticles (SPIONs), generate negative contrast in T2 weighted images. For clinical diagnostic applications, T1 agents are usually preferred because a number of natural sources (tissues with low signal intensity) also generate negative contrast, complicating the analysis of the MRI image. In fact, most FDA approved T2 agents were discontinued. Therefore the focus of this patent is on T1 agents.
A number of Gd-based CAs have been approved for clinical applications, such as ProHance® and Magnevist®, which are currently dominating the CA market. At least two problems exist with small Gd-based CAs. The relaxivity of Gd-based contrast agents could be higher, particularly at high magnetic fields e.g., at 3 T or higher. Further there is a problem of toxicity that results from free Gd(III) i.e., which escapes the chelating agent in certain patients with renal dysfunction. These two issues are interrelated.
Small Gd-based CAs have relatively low relaxivity (about 3-4 mM−1 s−1 at 1 T, 37° C.) and as a result, gram quantities are typically injected into a patient for generating an image with reasonable quality. In recent years, MRI scanners are moving to higher magnetic fields (mainly 3 T) in order to perform scans with improved signal/noise ratio, shorter acquisition times and better image resolution. Since the relaxivity of the commercially available Gd-based CAs decreases at higher magnetic fields, even larger quantities of CAs would be required for adequate contrast enhancement14. In addition, studies such as contrast-enhanced MR angiography and delayed contrast-enhanced myocardinal viability examinations require even higher CA doses and thus, further increase the risk of metal toxicity15. Although many in vitro studies have indicated that these Gd-based CAs are thermodynamically stable, the emergence and proliferation of nephrogenic systemic fibrosis (NSF) cases correlated to the usage of Gd-based CAs since the late 1990s suggests in vivo release and accumulation of toxic free Gd(III) in certain patients with renal dysfunction. Symptoms of NSF include severe skin induration, muscle restlessness and sometimes, physical disability. To address the severity of this safety issue, the FDA requires a “Black Box” warning label to be attached to all Gd-based CAs indicating possible adverse effects.
Various attempts have been made to improve the relaxivity mainly by increasing the size and thus rotational diffusion time (TR) of Gd-based CAs, based on the Solomon-Bloembergen-Morgan (SBM) theory. Common strategies include attachment of Gd CAs to proteins, dendrimers or polymers. Notably, at high magnetic fields (>1.5 T), the electron spin relaxation of Gd(III) dominates the inner-sphere relaxivity, therefore, the strategy of increasing TR becomes much less efficient to improve r for Gd-based CAs at high fields than low fields. Despite the moderate relaxivity increase, Gd toxicity is likely to persist in these macromolecular CAs16. In fact, because most of the conjugation chemistry involves the chelation sites, the Gd affinity will be lowered, contributing to higher risk of heavy metal leakage. In addition, the internal flexibility of these aliphatic dendrimers or polymers also contributed to the less than expected increase in relaxivity per Gd. Lastly, these large CAs are retained in the body for a longer period of time, leading to higher chance of Gd release than small Gd-based CAs.
There is thus a need to create a new generation of CAs that is more efficient and avoids the adverse effects of Gd toxicity. It will be desirable if the new generation of CAs are free of toxic heavy metals such as Gd(III) and exhibit high T1 relaxivity at high field.